The present invention relates to an apparatus for use in medical imaging comprising a readout circuit having an input for receiving a signal corresponding to a photon hitting a radiation detector. The application further relates to a detector assembly and to a combined PET-CT apparatus as well as to a method according to the preamble of claim 27.
Medical imaging techniques employing radiation detectors for detecting photons for example comprise Positron Emission Tomography (PET), X-ray computed tomography (CT), Single Photon Emission Computed Tomography (SPECT) and γ cameras. While the invention is not limited to any specific type of medical imaging, for illustrative purposes we shall explain the invention with specific reference to PET.
PET is a nuclear medicine tomographic imaging technique using γ rays. To conduct a scan, a short-lived radioactive tracer isotope which decays by emitting a positron and which has been chemically incorporated into a metabolically active molecule is injected into the patient, typically into the blood circulation. There is a waiting period while the metabolically active molecule becomes concentrated in the tissues of interest. Thereafter, the patient is placed in an imaging scanner.
As the radioisotope undergoes positive beta decay, it emits a positron which after a short travel encounters and annihilates with an electron, thereby producing a pair of γ-photons having an energy of 511 keV and usually traveling in opposite directions. The γ-photons are detected by a radiation detector typically comprised of a scintillator and an associated photodetector. The signal from the photodetector must then be readout by a suitable readout circuit. PET depends on simultaneous or coincident detection of the pair of γ-photons. Photons which do not arrive in pairs (i.e. within a few nanoseconds) are ignored.
Since most of the γ-photons are emitted at 180 degrees from one another during electron-positron annihilation, the source of radiation can be located along a straight line connecting the two radiation detector sites at which coinciding hits are detected. If the response of the radiation detector and the corresponding reader circuit is fast enough, it is moreover possible to calculate the location of the radiation source on said line from a difference in the arrival times. This method is called “time-of-flight” (TOF) measurement. However, this would require a time resolution of the measurement in the picosecond range which is currently difficult to achieve. Instead, typically statistics are collected from tens-of-thousands of coincidence events and equations for the total activity of each parcel of the tissue of interest along the above mentioned straight line is solved, such that a map of radioactivities can be constructed and outputted. Clearly, if time-of-flight information is available, the statistics needed for producing a high quality image would be smaller, and accordingly, a smaller dose of radioactive tracer isotope could be used which would be healthier for the patient. A prime goal of the above mentioned apparatus is therefore to provide a very fast readout channel allowing for measurements in the picosecond rather than nanosecond range.
FIG. 1 schematically shows an electronic readout channel according to prior art which is currently used in PET-systems. The readout channel of FIG. 1 comprises and avalanche photo diode (APD) array element 10 receiving light from a scintillator (not shown) upon a photon hitting the scintillator. The APD array element outputs an analog signal which is processed by a preamplifier 12 and a shaper 14 and is then inputted into an analog-to-digital (A/D) converter 16. The A/D converter 16 is clocked at high frequency (typically 40 to 100 MHz) and outputs a digital signal. The digital signal is then transferred to a FIFO buffer 18 and to a coincidence processor 20 for detecting coincidence of two γ-photons which can be assigned to a common electron-positron annihilation. The digital samples are processed by an algorithm implemented in a digital filter that determines the time-stamp and a signal of the amplitude. The architecture of FIG. 1 continuously samples analog signals and produces large data sets that due to their large size are generally processed off-line further down the image processing chain.
The prior art readout architecture of FIG. 1, however, has the following disadvantages.    1. Retrieving time information for sampled data requires additional hardware and algorithms that increase processing time and hardware costs.    2. The precision of the sampling of the signals depends on the signal shape. Shape variations lead to a decrease of precision in the evaluation of the timing of the corresponding event, in particular since PET events are randomly distributed over time and thus uncorrelated with the sampling clock cycle.    3. To obtain a higher timing precision and event counting rate, the clock frequency of the sampling A/D converter must increase which might lead to significant data overflow problems.    4. Modern PET scanners and in particular whole-body-PET scanners with a large field of view use increasingly more detector rings containing highly segmented arrays of scintillating crystals each being as small as 2×2 mm2, such that the number of required front-end electronic readout channels easily increases to more than one hundred thousand. For such a large number of channels the cost, power dissipation and complexity of signal processing using the readout architecture of FIG. 1 becomes extremely demanding.
As mentioned above, the invention is not limited to PET-imaging, and accordingly, in FIG. 2 a prior art readout architecture for X-ray detection is shown. In the shown front-end electronics of an X-ray CT-scanner system an APD-array element 10 collects light from a scintillator (not shown) hit by an X-ray photon and generates an analog current signal. The analog current signal is amplified by a current amplifier 22, and the amplified current is integrated by an integrator 24. The integral of the current received in a given time frame is proportional to the number of X-ray photons hitting the detector in said time frame. The integrated signal is then digitized by an A/D converter 26 and outputted to an image reconstruction processing logic. Note that in the context of this application, the term “logic” refers to any kind of hardware, software or combination thereof providing the respective logic function.
PET-scans and CT-scans are preferably performed simultaneously since they give in combination both anatomic and metabolic information. That is, modern PET-scanners are available with integrated high-end CT-scanners. Since the two types of scans can be performed sequentially without the patient having to move between the scans, the two sets of images are precisely registered such that areas of the PET imaging can be precisely correlated with the anatomy provided by the CT-images. Note however, that the readout technologies of FIGS. 1 and 2 are incompatible with one another such that it is not possible to perform a simultaneous PET-CT co-registration with a common detector head.
In other prior art PET systems the photon arrival time is detected using a constant fraction discriminator (CFD) that extracts the precise time-stamp of photon arrival for coincidence and time-of-flight measurement. CFDs are for example known for the detection of scintillator pulses having identical rise times which are much longer than the desired temporal resolution. Accordingly, it is not sufficient to use a simple threshold triggering which would introduce a dependence of signal peak height and trigger time, an effect which is called “time walk” and is described more in detail below. Instead, according to CFD triggering is not performed on a fixed threshold but on a constant fraction of the total peak height yielding trigger times independent from peak heights.
An example of a CFD circuit is shown in FIG. 3. A CFD circuit used for PET-applications is for example described in detail in “An Analog Signal Processing ASIC for a Small Animal LSO-APD PET Tomograph”, V. Ch. Spanoudaki, D. P. McElroy and S. I. Ziegler, Nuclear Instruments and Method in Physics Research, Section A, in press, such that a detailed explanation is omitted.
However, a readout circuit based on a CFD has two main disadvantages. Firstly, the integration of a fast CFD circuit in a monolithic CMOS that is now currently done for electronics system turns out to be difficult for timing precision of better than 100 ps. Also, the CFD circuits are quite expensive if produced in the necessary amount. Secondly, a CFD-based readout architecture only provides a time-stamp but not the signal amplitude, such that PET events cannot be reliably distinguished from background events and Compton events cannot be reconstructed.
Accordingly, it is an object of the current invention to provide a method and an apparatus for use in medical imaging that overcome the above mentioned drawbacks.